Opposite to conventional radiography wherein an intensifying luminescent phosphor screen directly emits luminescent rays and wherein said screen is not a storage medium, radiation image recording and reproducing techniques utilizing a radiation image storage panel, referred to as the stimulable phosphor screen, sheet or panel, are provided with a stimulable phosphor. With radiation image recording and reproducing techniques, the stimulable phosphor of the radiation image storage panel is caused to absorb radiation, which carries image information of an object or which has been radiated out from a sample. Said stimulable phosphor is exposed to stimulating rays, such as visible light or infrared rays, which causes the stimulable phosphor to emit light in proportion to the amount of energy stored thereon during its irradiation exposure.
The emitted fluorescent light is then photoelectrically detected in order to obtain an electric signal. The electric signal is further processed, and the processed electric signal is utilized for reproducing a visible image on a recording material. This way of working, making use of storage phosphor sheets or panels as an intermediate storage medium is also called “computed radiography”. This way of working is clearly differing from radiation detection by scintillator materials, directly emitting radiation upon X-ray irradiation as has e.g. been described in U.S. Pat. No. 5,171,996 for luminescent materials such as CsI, KI, RbI, CdS, ZnxCd1−xS, CdWO3, GaySe, Gd2O2S, La2O2S or PbOz.
As in radiography it is important to have excellent image quality for the radiologist to make an accurate evaluation of a patient's condition, important image quality aspects are image resolution and image signal-to-noise ratio.
For computed radiography signal-to-noise ratio depends on a number of factors.
First, the number of X-ray quanta absorbed by the storage phosphor screen is important. Signal-to-noise ratio will be proportional to the square-root of the number of absorbed quanta.
Second, the so-called fluorescence noise is important. This noise is caused by the fact that the number of photostimulated light (PSL) quanta detected for an absorbed X-ray quantum is small. Since a lot of the PSL light is lost in the detection process in computer radiography, fluorescence noise has an important contribution to the signal-to-noise ratio. It is important that, on the average, at least 1 photon is detected for every absorbed X-ray quantum. If this is not the case, many absorbed X-ray quanta will not contribute to the image and signal-to-noise ratio will be very poor.
This situation is most critical in mammography, where X-ray quanta are used with low energy. Softer X-ray will give rise to less PSL centres and, therefore, to less PSL photons than harder X-rays.
In computer radiography, a number of PSL centres are created by the absorbed X-ray quanta. Not all PSL centres are stimulated in the read-out process, because of the limited time available for pixel stimulation and because of the limited laser power available. In practice, only about 30% of the PSL centres is stimulated to give rise to a PSL photon. Since these photons are emitted and scattered in all directions, only 50% of the PSL photons are emitted at the top side of the storage phosphor screen, where they can be detected by the detection system. The emitted PSL photons are guided towards the detector by a light guide. This light guide may consist of an array of optical fibres, that forms rectangular detection area above the storage phosphor screen and has a circular cross-section at the detector side. This type of light guide has a numerical aperture of only 30%, which means that only 1 out of 3 of the emitted PSL photons is guided to the detector. In between the light guide and the detector a filter is placed, which stops the stimulation light reflected by the storage phosphor screen and transmits the PSL light emitted by the screen. This filter also has a small absorption and reflection of PSL light and transmits only ca. 75% of the PSL photons. In computer radiography a photomultiplier is commonly used to transform the PSL signal into an electrical signal. At 440 nm the photomultiplier has a quantum efficiency of ca. 20%. This means that only 1 out of 5 PSL quanta that reach the photomultiplier are detected.
In summary, for 1,000 PSL centres that are created in the storage phosphor screen only:1,000×0.3×0.5×0.3×0.75×0.2=6.75
PSL photons are detected. If it is required that every X-ray quantum gives rise to at least 1 detected PSL photon, therefore, the number of PSL centres created by an X-ray quantum should be sufficiently large. Or, conversely, the X-ray energy required to create a PSL-centre should be sufficiently small.
In mammography, a usual setting of the X-ray source is at 28 kVp. This leads to an X-ray spectrum, where the average energy of an X-ray quantum is of the order of 15 keV. For an X-ray quantum with this energy, in order to give rise to at least 1 detected PSL photon, the energy needed to create a PSL centre should be less than:15,000×6.75/1,000=100 eV.
Furtheron is well-known that fine detail visualisation, high-resolution high-contrast images are required for many X-ray medical imaging systems and particularly in mammography. The resolution of X-ray film/screen and digital mammography systems is currently limited to 20 line pairs/mm and 10 line pairs/mm, respectively. One of the key reasons for this limitation is associated with the phosphor particle size in the currently used X-ray screens. In particular, light scattering by the phosphor particles and their grain boundaries results in loss of spatial resolution and contrast in the image. In order to increase the resolution and contrast, scattering of the visible light must be decreased. Scattering can be decreased by reducing the phosphor particle size while maintaining the phosphor luminescence efficiency. Furthermore, the X-ray to light conversion efficiency, the quantum detection efficiency (e.g. the fraction of absorbed X-rays convertable to light emitted after stimulation) and the screen efficiency (e.g. the fraction of emitted light escaping from the screen after irradiation with stimulating rays) shouldn't be affected in a negative way by the reduction of the phosphor particle size. As a particular advantage the computed radiographic recording and reproducing techniques presented hereinbefore show a radiation image containing a large amount of information, obtainable with a markedly lower dose of radiation than in conventional radiography. Radiation image recording and reproducing techniques are thus efficient, particularly for direct radiography, such as the X-ray image recording for medical diagnosis.
For clinical diagnosis and routine screening of asymptomatic female population, screen-film mammography today still represents the state-of-the-art for early detection of breast cancer. However, screen-film mammography has limitations which reduce its effectiveness. Because of the extremely low differentiation in radiation absorption densities in the breast tissue, image contrast is inherently low. Film noise and scatter radiation further reduce contrast making detection of microcalcifications difficult in the displayed image. So e.g. film gradient must be balanced against the need for wider latitude.
Computed Radiography (CR) systems can be broadly categorized as primary digital and secondary digital systems. Primary digital systems imply direct conversion of the incident radiation on a sensor into usable electrical signals to form a digital image. Secondary digital systems, on the other hand, involve an intermediary step in the conversion of radiation to a digital image. For example, in digital fluoroscopy, image intensifiers are used for intermediary conversion of X-rays into a visible image which is then converted to a digital image using a video camera. The scintillator materials described in U.S. Pat. No. 5,171,996 e.g. do not require a scanning procedure to become read out as there is no storage of energy as in stimulable or storage phosphor materials. X-ray quanta are directly converted into emitted light, which further generates electrical charges in a photoconductive layer, thereby providing generation of a visible image. Scattering of directly emitted light is decisive for sharpness, opposite to sharpness of storage phosphors used in digital X-ray systems wherein scattering of stimulating light, as a source of photostimulation of stored energy is of crucial importance. Digital X-ray images generated in systems making use of photostimulated luminescence (PSL) plates, first store the virtual image as energy. In a second step, the stored energy is converted into electrical signals using a laser to scan the PSL plate to form a digital image.
Furthermore, various schemes using silicon photodiode arrays in scanning mode for CR systems have been employed. However, these photodiode arrays require intermediate phosphor screens to convert X-rays into visible light, because of the steep fall-off in quantum efficiency (sensitivity) of the arrays at energies above 10 keV.
The above described secondary digital systems have several disadvantages, including loss in image resolution. Recent technological advances have however made it possible to overcome these difficulties by allowing semiconductor X-ray detectors to be used to generate usable X-ray images. High quality semiconductor X-ray detectors have been known for many years, but these detectors require a very sensitive preamplifier to produce a signal suitable for use. With recent advances in high density analog complementary metal oxide semiconductor (CMOS) integrated circuit technology and high density interconnection between semiconductor chips, the integration of thousands of these detector elements with preamplifiers on a single hybrid integrated circuit called a sensor chip is now possible. A semiconductor detector having an absorbing layer located between X-rays from an object and X-ray semiconductor is sensors has e.g. further been disclosed in U.S. Pat. No. 4,905,265.
Although significant improvements of clinical image quality in order to eliminate the need for repeated exposures due to poor film image quality caused by factors as radiation scatter noise, fog, blurring, mottle and artifacts have meanwhile been realized in that digital radiographic techniques enable medicins to perform quantitative radiography through image digitization and allows them, by useful enhancement techniques, such as edge enhancement of microcalcifications and transmission of mammograms to remote sites over computer networks; advantageously reducing the absorbed radiation dose received by a patient by at least a factor of seven as compared to screen-film mammography, further facilitating mammography for routine screening of asymptomatic population in the 35 years and older age group by significantly enhancing the benefit to risk equation, furthermore significantly reducing the absorbed dose to the patients during a needle localization biopsy procedure which can require as many as 10 exposures. As has been set forth in U.S. Pat. No. 5,596,200 another advantage of that invention was that it provides improved storage and retrieval of image data through the use of standard magnetic or optical disk media instead of the photographic film, further providing a device which is capital cost competitive with current X-ray imaging systems and which reduces the cost in materials and processing time by eliminating photographic film and associated chemicals, dark rooms and other peripherals, as well as reducing technician's time for film processing. In determining the desired semiconductor materials therein one has to take into account aspects as ease of fabrication, X-ray absorption, and operating temperature. For mammographic applications, two alternative detector materials, silicon and gallium arsenide, are preferred. Silicon detectors are much easier to fabricate than GaAs detectors, however, the silicon X-ray photon quantum absorption coefficient is much lower than GaAs. For applications in a primary X-ray digital imaging system having X-ray energies greater than approximately 25 keV, sensor materials with much higher X-ray absorption properties are needed. Consequently, GaAs, cadmium telluride, CdZnTe, indium antimonide, and germanium are detector materials should be used at energies greater than 25 keV. The number of rows and columns of detectors and their length and separation can further be changed depending on the specific design requirements of the X-ray imaging system. For example the length of the row and the number of rows can be any desired value up to the limit of the mechanical scan. It is also contemplated to have the sensor chips placed in an array-like fashion.
Since the image generated is isomorphic to the matrix of digital numbers generated during the scan, it can be processed by a signal processing unit 60 with suitable software. For example, the signal-to-noise ratio of the signals can be improved through processing. Data from signal processing are advantageously stored and archived on standard magnetic or optical disk media instead of photographic film. Data from a storage station are then sent to an image processing unit in which a variety of processing operations can be performed on the image. For example, the image processing unit can perform the image manipulations of: (1) magnification; (2) contrast enhancement and windowing; (3) enhancing sharpness and edge gradients; (4) attaching gray or color scales to enhance image quality; and (5) image subtraction. Images generated by image processing can be displayed on a video display, a printer, on film, or sent via image transmission network, which can include satellites or computer networks to send image data from remote radiology laboratories to a centrally located radiologist for virtually real time image interpretation and diagnosis. Applications in other areas of clinical imaging are possible as e.g. for low dose, low cost applications in breast computed tomography (CT); use in intelligent software for computer aided diagnosis (CAD); the stereotactic computerized placement of biopsy needle; and radiation control, monitoring and non-invasive imaging systems for applications in nuclear medicine. An apparatus for imaging a patient's breast by scanning an imaging signal and a receiver across the patient's breast and then constructing a time-delay integration composite image based on the scan has been described in U.S. Pat. No. 5,526,394. Said receiver includes an array of radiation sensitive detector elements, wherein read out of the array is synchronized with the scanning motion of the receiver based on output from position encoder such that synchronisation is maintained despite scan drive variances. An assembly for allowing selection of an appropriate radiation filter based on particular imaging conditions has also been disclosed therein.
U.S. Pat. No. 6,300,640 specifically relates to a composite phosphor screen for detecting radiation, particularly X-rays, utilizing nanocrystalline sized phosphors (nanophosphors) disposed in extremely small channels (microchannels) etched in a substrate.
Further improvements in X-ray imaging, as will always be desired, more particularly in mammographic applications, making use of a particular detector, have been set out hereinafter.